Tissue engineering seeks to repair, replace or restore tissue function, typically by combining biomaterials and living cells. Control of polymer scaffold architecture is of fundamental importance in tissue engineering A wide variety of techniques for controlling the architecture of biomaterials are already available for relatively large feature sizes on the order of millimeters to centimeters. These include polymer extrusion, solution casting and particulate leaching, deposition of a polymer solution stream on a spinning mandrel and manipulation of sheets of polymer meshes. To achieve arbitrary three dimensional geometries, preformed sheets of biomaterial have been cut and laminated with a resolution of 0.8 mm. Such supports are useful for forming the macroscopic shape of the replacement tissue (i.e. an ear for cartilage tissue engineering) or for customizing tissues replacements for individualized patients (i.e. an eye socket for bone tissue engineering).
The properties of the tissue engineered construct emerge from the local response of the cells to their 3-dimensional microenvironment. Therefore it is of great importance to recreate biochemical and structural components of the in vivo cellular microenvironments when designing implantable tissue constructs. This microenvironment can be simulated by patterning of the matrix in which the cells are grown in or on, or by patterning the cells within the matrix. Traditional methods for controlling biomaterial scaffold architecture involve a number of methods, each with its own intrinsic limits related to the materials employed, its resolution or its costs. Injection molding against a microfabricated silicon template was utilized by Kapur et al (1996) with a resolution of 10 microns. A three dimensional printing technique developed by Griffith et al (1998) utilizes a polymer powder spread on a plate. The resolution of this method is dependent upon the polymer particle size where the typical features are on the order of 300 microns. These techniques are useful for forming complex tissues such as bone/cartilage composites for the knee and for optimizing microscale architecture to improve the function of the resultant tissue. For example, scaffold texture can alter cell migration, ingrowth, vascularization, and host integration. Microscale scaffold architecture can also modify the cellular responses such as growth and differentiation as has been shown on three-dimensional polymer meshes (e.g. U.S. Pat. No. 5,443,950).
Methods to prepare scaffolds with microscale structure that are more amenable to use with biodegradable polymers such as poly-DL-lactide-co-glycolide (PLGA) have also been developed. Material microstructure was first controlled by process parameters such as the choice of solvent in phase separation, doping with particulate leachants, gas foaming, woven fibers, and controlled ice crystal formation and subsequent freeze-drying to create pores; however, these scaffolds lack a well-defined organization that is found in most tissues in vivo (i.e. pores are randomly distributed rather than oriented and organized in functional units). Similarly, microtubular scaffolds (Ma and Zhang, 2001); 3-dimensional micropatterned scaffolds using UV polymerization (Ward et al., 2001) produce scaffolds with arbitrary architectures. The use of soft lithography methods using biopolymers such as poly(dimethylsiloxane) (PDMS) allows for the production of high resolution 2-dimensional scaffolds that may be assembled into higher order structures (Vozzi et al., 2003). However, the method is cumbersome for the production of 3-dimensional structures as the maximum thickness possible for each scaffold layer is about 30 microns.
None of the methods discussed above allow for the generation of more, complex cellular tissue constructs in which cells can be placed in specified 3-dimensional configurations throughout a thick construct. For example, biomaterial scaffolds must be seeded with cells with the help of gravity, centrifugal forces or convective flow (Yang, et al., 2001). Alternatively, cells can be recruited to the graft by the use of growth factors and chemokines (Badylak et al., 2001). Many of the techniques described above require processing conditions such as heating and polymer grinding that may be limiting for the inclusion of bioactive moieties and preclude the formation of scaffolds in which the cells are cast in the scaffold. In the methods discussed above, the cells are patterned based on the physical structure of the surfaces of the scaffold. None of these methods enable formation of a thick tissue construct that is populated with living cells.
A number of methods have been developed for the generation of essentially two dimensional cell arrays. Cells may be positioned in or on a substrate using mechanical methods such as pipette/syringe placement (e.g. Landers et al., 2002), stenciling (Folch et al., 2002) and microfluidic delivery (Folch et al., 1999) and by optical methods such as optical tweezers (reviewed by Ashkin, 1991) and laser-guided writing (Odde et al., 2000). Electromagnetic forces such as elelctrophoretic, dielectrophoretic (DEP) (Matsue et al., 1997) and magnetic attraction/repulsion, may also be used for bulk patterning of cells in a two dimensional space. However, in the absence of an appropriately adhesive substrate, the cells return to a random state after the electromagnetic forces are removed. Moreover, patterning of cells in 2 dimensions is of limited utility for long term maintenance of differentiated cells. Cells sitting on a 2-dimensional surface often spread and lose function depending on both the cell types to be cultured and the size of the regions of each cell type. Chondrocytes, for example, have a rounded morphology in vivo and cease normal biosynthetic activity upon spreading into a flattened shape on a 2-dimensional surface. Methods of co-culturing have been used with some success to maintain hepatocytes in a differentiated state. However, these methods are cumbersome and are typically not useful for long term maintenance of cells in culture. Bhatia et al. (U.S. Pat. Nos. 6,130,479 and 6,133,030) teach methods of patterning various cell specific adhesion molecules (e.g. collagen) on glass slides using photoresist, a UV-sensitive polymer, and photolithographic techniques. Maintenance of a differentiated, state in hepatocytes was dependent on the ratio of surface to perimeter of the islands of hepatocytes within the non-parenchymal cells. High resolution patterns (10 microns) can be generated for the co-culture of cells with, different adhesion properties, however, the method is still limited to only two dimensional patterns and great care must be taken to maintain the cells in the desired state.
Methods have been developed for patterning of cells within a matrix by patterning molecules within the cell seeded microenvironment, by magnetic orientation of fibrin in gels (Dubey, 2001) or by stretching an underlying support (Vandenburgh et al, 1982). However, these methods orient all cells in a volume in a constant orientation, either linear or radial. Specific microscale patterning and positioning is not possible. Cells may be encapsulated in a 3 dimensional support, such as a hydrogel, but it is not possible to achieve specific cell patterns or orientations. These methods do not allow for the precisely placement of cells in a matrix; rather a population of cells is generally oriented or confined to a specific volume much larger than the cell.
Hydrogels are becoming an increasingly popular material for tissue engineering because their high water content and mechanical properties resemble those of tissues of the body. In addition, many of the hydrogels can be formed in the presence of cells by photopolymerization, which allows homogeneous suspensions of cells throughout the gel. Poly(ethylene glycol) (PEG)-based hydrogels are of particular interest because of their biocompatibility, hydrophilicity and the ability to be customized by changing the chain length or chemically adding biological molecules (Peppas et al., 2000, incorporated herein by reference). These types of hydrogels have been used to homogenously immobilize various cell types including chondrocytes (Elisseeff et al., 2000) vascular smooth muscle cells (Mann et al., 2001) and fibroblasts (Gobin and West, 2002; Hem and Hubbell, 1998) that can attach, grow and produce matrix.
One property of these hydrogel systems that has not yet been exploited is the use of the photopolymerization step to form structural 3-dimensional hydrogel features containing cells. Elsewhere in non-biological systems, the fundamentals of photolithography have indeed been applied to PEG-based hydrogel systems to create hydrogel valves within flow systems by controlling regions of photopolymerization using a mask (Beebe et al., 2000). This process would not be amenable to the incorporation of living cells due to the harsh chemical conditions and the high density polymers with short polymer chains used to obtain high resolution structures.
One goal of tissue engineering is the production of artificial tissues or organs for transplant (e.g. cartilage). Cartilage, for example, is an avascular tissue with, little or no capacity for effective repair following traumatic injury, due to a limited cell population near the injury that is encased in a dense matrix. Surgical transplantation of cartilage tissue improves patient function in the short-term but suffers from limited donor supply and donor site morbidity. Therefore, tissue engineering approaches are in development to address the tissue sourcing problem by forming cartilaginous constructs from minor biopsies. However, producing such constructs with appropriate biological and mechanical properties requires an understanding of the complex cellular architecture and, potentially, a method for controlling the cellular architecture.
Another goal of tissue engineering is to develop organ support systems such as an artificial liver apparatus, similar to a kidney dialysis apparatus, for hepatic support in individuals waiting for liver transplant. A number of artificial liver devices have been developed (e.g. Naughton, U.S. Pat. No. 5,827,729, incorporated herein by reference), most of which require viable, differentiated hepatic cells in order to function. The liver is a more complex organ than the kidney which is predominantly responsible for salt balance and filtering of molecules based on size. The liver is responsible for detoxification of xenobiotics and hormones, energy metabolism, production of plasma proteins, and production of bile, rather than the simple filtering, of the blood. Furthermore, the factors that lead to hepatic coma in patients suffering from liver failure have not been identified. Therefore, sustenance of a patient in liver failure with a device that lacks hepatic cells is unlikely.
The development of a method to allow for the growth and maintenance of primary hepatocytes would be useful in developing a better understanding of drug metabolism and interactions. A simulated liver could be used desirable for the testing of drugs, both alone in the process of drug development, and to better understand drug interactions (Hodgson, 2001). Initial drug testing is typically performed on cells in culture to facilitate high throughput screening. However, compounds ingested by a patient must have desirable ADMET (Absorption, Distribution, Metabolism, Elimination and Toxicity) properties in order to be successful as a drug. Such tests can be performed in animals, however there are a number of drawbacks including expense, variation between species, and growing disfavor of the use of animals in research by the general public. However, the maintenance of a culture of differentiated hepatocytes is non-trivial. Systems to study the effects of liver metabolism include the use of liver slices, immortalized cell lines and isolated liver enzymes have been developed. Each system is limited by various factors including variability between species, phylogenetic drift of cell lines and possible inaccuracies of using liver enzymes in isolation. The development of a method to allow for the growth and maintenance of primary hepatocytes would be useful in developing a better understanding of drug metabolism and interactions.